Thermal coagulation therapy may be used for the treatment of localized diseased tissue, e.g., tumors, in a diseased organ or body. Generally, a target volume of tissue is sufficiently heated to achieve a therapeutic effect, such as thermal coagulation. Tissue thermal coagulation depends on a number of factors, and temperatures in the range of 55-60° C. are generally considered sufficient to provide enough energy to cause such coagulation. Cell death results from heating to these temperatures, and a region of irreversible thermal damage can be observed with imaging following the treatment. In addition, heating can be produced from minimally-invasive applicators, eliminating the need for open surgery, and potentially reducing recovery time and morbidity for patients. This approach has been used with some success in the treatment of isolated primary liver cancers and colorectal metastases for patients otherwise ineligible for surgery.
Interstitial thermal therapy is currently practiced by inserting heating applicators directly into a target site within an organ. Several energy sources have been integrated into interstitial heating applicators, including lasers, ultrasound, microwave, and radiofrequency energy. Preferably, interstitial thermal therapy delivers sufficient thermal energy to coagulate an entire target volume, while avoiding undesirable thermal damage to normal tissue. This strategy is referred to as “conformal thermal therapy.” One limitation of present interstitial thermal therapy technology is the inability to control or adjust the three-dimensional pattern of energy deposition dynamically during a treatment. Most current applicators act as point or line sources of energy resulting in highly symmetric patterns of energy deposition in tissue. This makes it difficult to treat targets with complex geometry accurately, and does not take full advantage of the imaging information available with imaging technology such as magnetic resonance imaging (MRI).
One application of interstitial heating is transurethral prostate thermal therapy, which selectively destroys diseased prostate tissue using a device located within the prostatic urethra, and preserves adjacent normal tissues such as the rectal and bladder walls. Disease targets include prostate cancer and benign prostatic hyperplasia (BPH). Current transurethral thermal therapy technologies are incapable of producing a thermal treatment (cell death) pattern that conforms accurately to the geometry of the prostate gland or to selected regions of the prostate gland or other organ or tissue undergoing treatment.
In conformal prostate thermal therapy applications, it is often desirable to implement some form of quantitative temperature monitoring for feedback during treatment to ensure accurate delivery of energy to the prostate gland. Temperature monitoring of treated (or heated) tissue regions can be accomplished in several ways. These include direct measurements as well as indirect measurements of the spatial and/or temporal thermal field in the treatment region.
Magnetic resonance imaging (“MRI”) has been used to non-invasively measure spatial heating patterns in tissue. Several MRI techniques are available to measure the temperature distribution in tissue. These techniques are possible because of the temperature dependence of various nuclear magnetic resonance (“NMR”) biophysical parameters such as T1, T2, diffusion, magnetization, and proton resonant frequency. The most commonly used technique in MRI-guided thermal therapy is the proton resonant frequency (“PRF”) shift technique, which exploits the direct proportionality between the resonant frequency of water protons and temperature. A common technique to measure this effect employs the subtraction of a baseline phase image obtained prior to heating from a phase image obtained during heating to measure the change in phase resulting from local temperature elevations. The change in phase can then be related to the change in temperature through the expression,
      Δ    ⁢                  ⁢          T      ⁡              (                  x          ,          y                )              =            Δ      ⁢                          ⁢              Φ        t                    α      ·      γ      ·              B        o            ·      TE      where ΔT is the temperature change between two images, ΔΦ is the phase change between the same two images, α is the proton resonant frequency shift coefficient (typically −0.01 ppm/° C.), γ is the gyromagnetic ratio, Bo is the strength of the main magnetic field (T), and TE is the echo time of the imaging sequence used to acquire the two images.
A phase subtraction technique is required due to large static background phase variations in MR images due to inhomogeneities in the main magnetic field, which are much larger than the phase changes due to temperature. Unfortunately, MR thermometry using the phase subtraction technique is limited in its accuracy by a number of issues including drift in the main magnetic field, inter-image motion, susceptibility-induced temperature artifacts, inaccurate baseline estimates, and relative motion of the detector and the measured subject.